In the field of nuclear medical imaging technology, a subject, e.g. an oncology patient or an animal used in an experiment, can be scanned by detecting radiation emanating from the subject. For example, in a so-called PET scan, a short-lived radioisotope, which decays by emitting a positron, is injected usually into the blood circulation of a living subject. After the metabolically active molecule becomes concentrated in tissues of interest, the research subject or patient is placed in the imaging scanner. The most commonly-used metabolically active molecule for this purpose is 18F-fluorodeoxyglucose (FDG), a sugar, which has a half life of 110 minutes.
As the radioisotope undergoes positron emission decay, it emits a positron, the antimatter counterpart of an electron. After traveling up to a few millimeters, the positron encounters and annihilates with an electron, producing a pair of gamma photons moving in almost opposite directions. These are detected when they reach one of a plurality of scintillation crystals in the scanning device, creating a burst of light detected by an array of photosensors.
Radiation emanating from the subject can be detected in, for example, radiation detector ring assembly 13 illustrated in FIG. 1. At a more granular level, specific radiation events can be detected at detector block 16 comprising an array of radiation sensors, such as plurality of scintillators and associated photosensors 11n, such as photomultiplier tubes (PMTs), avalanche photodiodes (APDs), or silicon photomultipliers (SiPMs). In the case of a PET scan, scintillators can be arranged in a ring 13.
Generally a plurality of sensors, e.g., photosensors 11, can be arranged in a matrix and assigned to detect the light of a single scintillator as shown in detector block 16 in FIG. 1. Detector block 16 can be associated with a single scintillation crystal 13 or can be, as shown, a matrix of scintillator crystals that is coupled to the photosensors 11 1 . . . 11n usually via a light guide. A plurality of detector blocks 16 can be axially arranged adjacent to one another, in a slot, in a line relative to the center of ring 13. To be able to increase the resolution of the system without the high costs of 1:1 coupling, the number of photosensors 11 per block is generally significantly lower than the number of scintillation crystals 13. For example, a detector block may have a plurality of radiation sensors, such as photosensors 11 with, for example, 4, 9 or 16 photosensors 11 arranged in a 2×2, 3×3, or 4×4 matrix behind an array of scintillation crystals 13. Other arrangements with more or fewer photosensors 11 are possible. Thus, scintillation event localization can be determined or interpolated by such a detector block by processing the associated photosensor signals. This can be done by analog filtering, integration, and multiplication of weighted combinations of the photosensor signals or by using digital algorithms that use discrete time sample points of signals obtained directly from the photosensors 11. The PET technique depends on scintillation event detection of the pair of gamma photons.
FIG. 1 illustrates a block diagram of the typical architecture of a detector block 16 and associated analog-to-digital-converters 14-14n in a conventional system. Each matrix of photosensors 11 produces a plurality of signals that can be processed to generate an image from a plurality of scintillation events that are detected by a photosensor 11. To determine the location of a detected annihilation, the system needs to accurately measure the timing and energy of both detected photons. Consequently a high amount of data has to be produced by the respective measurement circuits.
For example, as shown on the right side of FIG. 1, each scintillator has an associated matrix of detector blocks, such as photosensors 111 . . . 11n, which, in this example are PMTs. Each signal of each PMT 111 . . . 11n is first amplified by, for example, associated preamplifiers/buffers 121 . . . 12n. The output signal of preamplifier/buffers 121 . . . 12n can then be converted concurrently into discrete-time digital signals by associated analog-to-digital converters (ADC) 141 . . . 14n. A sampling clock for each ADC be can provided at terminal 15. In this example, this digital processing architecture uses n independent ADC signals with peripheral circuitry to concurrently sample each of n photosensor signals per block. This can increase the costs of a detector block.
FIG. 2 illustrates a detector block 20 comprising an 8×8 array of scintillation crystals; for example, each crystal can be 4 mm×4 mm×20 mm. Photosensors 11 can be included behind the scintillation crystals to detect light emitted due to scintillation events.
Not all radiation emanating from a subject is detected by scanner 10. Radiation can be emitted outside of the field of view of scanner 10, or radiation can scatter. For example, Compton scatter, which can occur when a photon collides with an electron, thereby transferring energy to the electron. The collision can cause the photon to deviate from its original path and cause a loss of energy. This collision typically occurs within the subject or in, for example, a scintillation crystal. Due to Compton scattering, events that would otherwise have been detected may be missed.
The probability that a 511 keV gamma ray be detected is a function of the material composition of the detector block, its size, and its density. For LSO, the probability that the first interaction of the 511 keV gamma ray is a Compton scatter is on the order of 68%, and for short, narrow pixels, the fraction of Compton scatter exiting the pixel can be quite significant.